Scintillation crystals have conventionally been used in non-invasive medical diagnostic techniques that utilize radiation-emitting materials. One such technique is positron emission tomography (PET), which provides in-vivo, functional information about the molecular biochemistry of a given radio-labeled compound (tracer) introduced into a live subject. The radio-label is a positron emitter, which results in high energy photon emissions when positrons annihilate with electrons in body tissue. The tomographic imaging is possible through detection and localization of the many associated highly energetic photons emitted.
In conventional PET, the photons are absorbed in a scintillation crystal, which gives off a flash of light. The light is collected by a photodetector, which detects and converts the light into electric charge that is amplified. The result is a robust electric signal with an amplitude that represents the energy of the incoming photon, a location that indicates where the energetic photon came from within the imaging subject, and a time stamp that signifies when the event occurred. For high spatial resolution imaging, which will allow one to see very minute structures, conventional PET relies on very accurate localization of the energetic photon emissions. This means that the scintillation detector must have very fine position resolution of the entering photons. However, to efficiently absorb the incoming photons, the crystal must also be relatively thick. Efficient absorption of incoming photons is important to allow for high count sensitivity, which translates into good image quality. Further, the signals that are created should be as robust as possible.
The state of the art was advanced by the invention described in U.S. Pat. No. 6,114,703 to Levin et al. The '703 patent provided an efficient method and devices for collection, and made the large surfaces of long and narrow scintillation crystals available for detection. The '703 patent disclosed methods and devices that replaced the bulky and expensive photomultipliers (PMTs) by utilizing semiconductor photodetectors, applying such semiconductor photodiodes directly to surfaces of the scintillation crystals, including at least one large surface of the scintillation crystal. The device of the '703 patent improved the amount of light measured from a scintillation event, while maintaining high spatial resolution offered by long and narrow scintillation crystals. The '703 patent also improved upon the single sheet style conventional devices that receive radiation in the large face of the crystal sheet by eliminating the coupling losses associated with the optical interfaces between the crystal and PMT and replacing the PMT of the conventional devices with directly deposited semiconductor photodiodes.
An overriding goal in radiation imaging is to obtain reconstructed images of very high spatial resolution. Spatial resolution improvements in reconstructed images have come most often from reductions in the size and increases in the number of scintillation crystals. Detection sensitivity, though, is another limiting factor. The '703 patent was directed to improvements in the detection sensitivity. To maintain high detection sensitivity and good image quality, the challenges were to develop a finely pixellated scintillation crystal array with both high detection efficiency and high light collection. High detection efficiency means the crystals must be relatively long, tightly packed, and cover a relatively large axial field-of-view (FOV). High spatial resolution means that the crystals are very narrow.
A difficulty with designs having small scintillation crystals for high resolution is that manufacturing is a significant challenge. It is costly and complex to handle many minute crystal elements and align them with corresponding photodetector elements. Slight misalignments might reduce light collection efficiency. A shortcoming with conventional crystal sheet devices for PET is that the sheet must be thin so that it produces a relatively narrow beam of light onto the photodetector plane. Thus, crystal sheet detectors (e.g. coincidence gamma ray cameras) that have been used in PET suffer from low efficiency for stopping the high energy photons.
A prior application Ser. No. 10/664,768, now published as US-2004-0124360-A1, filed Sep. 17, 2003 (the '768 application) provides additional background for the present invention. The '768 application discloses, among other things, scintillation crystal sheets arranged in stacks parallel to each other. Semiconductor photodetector positional detectors read light from large faces of the scintillation crystal sheets to detect interactions in the scintillation crystal sheets and independently provide positional information concerning the interactions relative to two axes.
A preferred embodiment in the '768 application includes an array of scintillation crystal sheets arranged in a device such that radiation is incident upon small end faces of the sheets (“end face geometry” or “edge-on” geometry), and is fully described in the '768 application. Semiconductor photodiodes read light from large faces of the crystal sheets. The semiconductor photodiodes in the '768 application may be pixellated, meaning that the semiconductor diodes provide both detection of photons generated in the scintillation crystals and positional information about a detection, or may be, one large pixel with positioning capability within that pixel. In another preferred embodiment of the '768 application, radiation is incident on a large face of scintillation crystals (“large face geometry” or “face-on geometry”).